This application claims the benefit of a priority under 35 U.S.C. 119 to Great Britain Patent Application No. 0423325.0 filed Oct. 20, 2004, the entire contents of which are hereby incorporated by reference.
This invention generally relates to a radio frequency (RF) and gradient coil arrangement for a magnetic resonance (MR) imaging system or scanner. More particularly, the invention pertains to an arrangement wherein the RF coil of the MR scanner is selectively mounted in relation to other components of the scanner, to substantially increase RF coil performance, by maximizing separation therebetween while sufficiently cooling the gradient primary coil.
As is well known, a MR imaging system or scanner commonly includes a cryostat, which contains a powerful superconductive main magnet positioned around a main magnet bore. The superconductive magnet is maintained at an extremely cold temperature and produces a strong static magnetic field, or B0 field, within the bore, the B0 field being directed along the bore axis. Other essential components of the MR system include the RF coil, or RF antenna, and the gradient coil assembly, which comprises a hollow cylindrical structure. The RF coil may be operated in a transmit mode, to generate MR signals in an imaging subject, or may be operated in a receive mode to detect the MR signals. The gradient coil assembly comprises one or more cylindrical coil forms, as well as a set of gradient coils supported thereby, to produce the X-, Y-, and Z-gradient magnetic fields. These fields are required to spatially encode MR data. Typically, the gradient coil assembly is positioned within the main magnet bore.
In the past, it has been common practice to support the RF coil within the main magnet bore by attaching it to a further essential MR system component comprising an inner cylindrical form. The inner form comprises a tubular member which is inserted through the gradient coil assembly, in coaxial relationship therewith. The interior region of the inner tubular member generally comprises the patient bore or imaging volume of the associated MR system, that is, the volume which is disposed to receive a patient, and in which MR signals are generated and detected. The ends of the inner tubular member are attached to the cryostat, by means of end caps or the like, so that the tubular member is supported thereby. Typically, the RF coil is placed around the outside diameter of the inner tubular member, in close adjacent relationship, and supported or carried thereby. The tubular member is made of a non-electrically conductive material, so that it does not impede RF performance within the imaging volume.
As is well known, an MR system requires a RF coil to generate the B1 field, necessary to produce MR data while the set of three gradient coils are required to spatially encode the MR data. Both the RF and gradient coils are positioned around the MR magnet bore, which receives the patient or other imaging subject. The gradient coils are typically positioned outside the RF coil so that fields generated by the gradient coils must extend past the RF coil in order to reach the magnet bore.
In a typical coil assembly, the radio frequency coil is located within the gradient coils with a relatively small spacing therebetween. The close physical proximity of these different coils results in a significant amount of the energy from the RF excitation field being lost due to impingement upon the gradient field coil structure. This loss shows up as a damping of the quality factor (Q) of the contained radio frequency coil which, in turn, degrades the normally attainable signal-to-noise ratio of the imaging device. Accordingly some type of RF shielding is usually placed between the radio frequency and gradient coils to preserve the Q of the former coil and consequently the signal-to-noise ratio of the system.
If coupling occurs between the RF and gradient coils, the gradient coils will absorb or “suck away” RF coil energy. This causes the Q, or quality factor, of the RF coil as well as the signal-to-noise ratio thereof, to be significantly reduced. To prevent such absorption of RF coil power or energy, it has become common practice to place a RF shield between the RF coil and the gradient coils. From the standpoint of preventing RF coil power absorption, an ideal shield would be a cylinder of solid copper, or like conductive material, positioned between the RF and gradient coils. The RF field would induce image currents in the solid copper cylinder, which would serve to reflect the RF field so that it could not interact with the gradient coils. However, such a solid shield would also seriously attenuate gradient fields extending into the MR magnet bore.
In prior art arrangements, the RF coil and RF shield are located between the patient enclosure and gradient primary coils. To increase RF coil performance, the separation between the RF coil and RF shield should be increased while maintaining the RF shield disposed inside the gradient coils. To increase gradient coil performance, the gradient primary coils should be located as near to the patient enclosure as possible. The inside surface of the gradient primary coils needs to be cooled to prevent the coils from overheating and to maintain the patient enclosure temperature within regulatory limits. However, this cooling layer takes up space between the RF shield and primary gradient coils. Positioning the RF shield between the cooling layer and gradient primary coil decreases the cooling performance because the gap between the cooling layer and gradient coil is increased.
Accordingly, there is a desire to maximize cooling by minimizing the gap between the cooling layer and gradient coil, while maximizing RF performance by locating the RF shield next to the gradient primary coil.